The lifetime of enzyme biosensors is highly dependent upon the stability and the amount of immobilized enzyme inside the sensing element that is generally composed of an enzyme-containing multilayer membrane. How to most effectively prevent the loss of enzyme from the sensing element when a biosensor is used repeatedly or continuously remains a difficult problem. Most commercially available biosensors used in the food industry, environmental monitoring and blood analysis are based on replaceable sensing elements. Therefore, the lifetime of such a biosensor may be extended limitlessly by replacing the sensing element. Unfortunately, the replacement of the sensing element is not feasible for miniaturized biosensors, especially for implantable biosensors.
Miniaturized biosensors are well suited for continuous and in situ field monitoring. They have an obvious significance for medical applications. This is because a small biosensor needs only a small amount of sample for analysis and results in minimal surgical trauma when implanted for in vivo analysis. The most important progress made by miniaturized biosensors is to lead to the application of inexpensive, disposable maintenance-free biosensors in clinical and biomedical areas. Typical examples are i-STAT silicon-based single-use biosensors, FreeStyle™, TheraSense electrochemical test strips for blood glucose. Single-use biosensors are produced in mass by thin film fabrication technology and have a very good sensor-sensor consistency. The same technology is also used to manufacture miniaturized biosensors with a relatively long lifetime for critical care analysis, e.g. GEM Premier™. The lifetime of such a miniaturized biosensor is typically in the range of 2-8 weeks in continuous use with minimal maintenance. The lifetime is even less if it is used for implantation. A flexible sensor designed for subcutaneous in vivo amperometric monitoring of glucose was disclosed in U.S. Pat. No. 6,514,718 to Heller et al. The typical three or four-layered sensing element was employed in this sensor. The layers included an enzyme layer, a glucose flux limiting layer, a horseradish-peroxidase-based interference-eliminating layer and a biocompatible layer. The sensing element is formed within a recess upon the tip of a polymide insulated gold wire. The location of the sensing element requires stringent preparation techniques and allows for only a limited space for enzyme loading. To improve the response characteristics of an implantable glucose sensor in a low oxygen environment, a new kind of sensing element containing perfluorocarbon emulsion has recently been introduced (Analy Chim Acta 411 (2000) 187-192; see also U.S. Pat. No. 6,343,225 to Clark, Jr.). However, the application of this technology in the field of miniaturized biosensors may be limited because the presence of perfluorocarbon, which inevitably causes phase separation in sensing membranes, thereby impairing response stability as well as shortening the lifetime of the sensor.
The long-term stability of the sensing element is the development bottleneck of implantable biosensors. Various efforts have been made on improving the lifetime of implantable glucose biosensors over the past 30 years, but essential breakthroughs have materialized. The remaining difficulties with long-term implantable biosensors mainly originate from the strict requirements of implantation applications. To minimize surgical injury and discomfort caused by implantation, the probe size (diameter) is generally required not to exceed 1 millimeter. Thus, many of the mature fabrication technologies used for conventional biosensors are not applicable for implantable biosensors. In addition, the measurement environment for implanted biosensors is much more aggressive than blood or solutions. Therefore, an implantable glucose biosensor can only last for a few days in the body. So far, such biosensors are mainly used to provide a continuous real-time glucose variation profile for diagnostic and treatment optimization. Only a couple of continuous glucose monitoring devices have been successfully used for short-term glucose measurements in the skin by diabetes patients (e.g. Mini GGMS® and GlucoWatch®).
The function failure of an enzyme biosensor is mainly a result of the following three factors:
The loss of enzyme activity and/or the enzyme itself, including the separation of enzyme layer from the electrode surface;
Degradation of the polymer membrane (This can result from decomposition and physical damage of outer membrane, leakage of the sealing interface between membrane and electrode); and
Biological contamination of the outer membrane.
Using an adequate enzyme methodology is extremely important for the construction of a long-term, stable, miniaturized biosensor. Free enzyme loses more than 80% of its activity after one week in a solution, while immobilized enzyme can maintain catalytic activity for months and even years. Methods used to immobilize enzymes in biosensors include (1) adsorption; (2) physical entrapment; (3) chemical cross-linking; (4) covalent coupling; and (5) co-deposition. Among the various methods, the cross-linking method is most frequently used because it has the advantage of the covalent bonding, high enzyme loading and small loss of enzyme activity. Cross-linking agents, e.g. glutaraldehyde, bisisocyanate, bisdiazobenidine and chromium acetate are often used together with functionally inert proteins such as bovine serum albumin (BSA) and gelatin. The enzyme can also be cross-linked with epoxy resins as described in Biosensor and Bioelectronics, 11(8), 735-742 (1996). The disadvantage of chemical cross-linking is that the resulting cross-linking tends to separate from the electrode surface after wetting due to the poor adhesion of the gel layer to the electrode. Conversely, an electrodeposited enzyme layer has a stronger bond with the electrode surface, but the enzyme loading is very small. Electrochemical formation of non-conductive polymer is self-limiting, so the thickness generally does not exceed 100 nm.
The polymer membrane plays the role of enzyme layer protection and provides a diffusion-limiting barrier. It is readily understood that the property of the polymer can influence the biosensor's long-term performance. In many cases, the polymer membrane consists of several different polymer layers to obtain certain specific properties such as optimal response stability, good mechanical strength, high diffusion resistance for unexpected species and macromolecules and biocompatibility [see for instance U.S. Pat. Nos. 6,514,718, 5,773,270 and 4,418,148]. Unfortunately, true multi-layered polymer membranes cannot be realistically achieved by coating due to problems of inter-solubility and phase separation of polymers. Furthermore, multilayered membranes result in excess thickness which rapidly increases the response time of the sensor.
For these reasons, employing a layer of polymer membrane is an attractive option for overcoming the aforementioned limitations. Various polymer and copolymer materials have been investigated for biosensors, e.g. Nafion®, polydimethylsiloxane (PDMS), polycarbonate (PC), polyurethane (PU), poly(vinyl chloride) (PVC), cellulose acetate (CA), tetrafluoroethylene (Bull. Krean Chem. Soc., 24(4), 2003, 514-516), perfluorocarbon polymer (T. Matsumoto et al. Biosensors & Bioelectronics 16 (2001) 271-276), UV-curable epoxy acrylates (E.P. Pat. No. 1,292,823 (2003)) vinyl polymer with a siloxane region and epoxy group (J.P. Pat. No. 03,024,757) and polyallylamine-polyaziridine (U.S. Pat. No. 6,514,718). The selection of suitable polymers for sensing elements has a profound influence on the long-term properties of the resulting biosensor. The polymer layer should result in minimal adhesion of proteins and cells. Studies have addressed improving the physical nature of polymer membranes using various additives such as plasticizers according to the modification method of commercial films. For instance, using a plasticizer may interpose itself between the polymer chains and interact with the forces held together by extending and softening the polymer matrix and possibly leading to some improvement in brittleness, flexibility and strength, as well as increasing the adhesiveness of the film with other surfaces or layers.
For implantation applications, the additives must not cause toxicity as they leach out of the membrane. Furthermore, the polymer must have excellent biocompatibility. The fouling of biosensors commonly results from tissue reactions in the microenvironment around the membrane and microorganism deposition on it. Tissue reactions cause an increase in the diffusion resistance of the analyte and oxygen or can result in aberrant analyte distribution, and even cause failure of response. Improving the surface adsorption characteristic of the outermost membrane may inhibit the fouling tendency of a biosensor.